Embodiments of the invention relate generally to x-ray imaging devices and, more particularly, to an X-ray tube having an improved control of electron beam emission, and thus, of X-ray generation.
X-ray systems typically include an X-ray source or tube, a detector, and a support structure for the X-ray tube and the detector. In operation, an imaging table, on which the object is positioned, is located between the X-ray tube and the detector. The X-ray tube typically emits radiation, such as X-rays, toward the object. The radiation typically passes through the object on the imaging table and impinges on the detector. As radiation passes through the object, internal structures of the object cause spatial variations in the radiation received at the detector. The data acquisition system reads the signals received in the detector, and the system translates the radiation variations into an image, which may be used to evaluate the internal structure of the object. One skilled in the art will recognize that the object may include, but is not limited to, a patient or subject in a medical imaging procedure and an inanimate object as in, for instance, a package in an X-ray scanner or computed tomography (CT) package scanner. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged with X-rays.
Typically, in an imaging system such as a computed tomography (CT) imaging system, an X-ray source emits a fan-shaped or cone-shaped beam toward an object, such as a patient or a piece of luggage. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the X-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis, which ultimately are used to produce an image.
X-ray detectors typically include a collimator for collimating X-ray beams received at the detector, scintillator adjacent the collimator for converting X-rays to light energy, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts X-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
X-ray tubes typically include an anode structure or target for the purpose of distributing heat generated at a focal spot. An X-ray tube cathode provides an electron beam from an emitter that is accelerated using a high voltage applied across a cathode-to-anode vacuum gap to produce X-rays upon impact with the anode. The area where the electron beam impacts the anode is often referred to as the focal spot. Typically, the cathode includes one or more filaments positioned within a cup for emitting electrons as a beam to create a high-power large focal spot or a high-resolution small focal spot, as examples. Imaging applications may be designed that include selecting either a small or a large focal spot having a particular shape, depending on the application.
In the following paragraphs a more detailed description will be provided of a CT system that implements the improved method of electron beam control that is the subject of this invention. However, it is to be understood that the invention is not limited to CT systems but can be applied to all types of imaging systems that include one or more X-ray tubes.
A CT imaging system may include an energy discriminating, multi energy, or dual energy capability. Techniques to obtain the measurements may include scanning with two distinctive energy spectra and detecting photon energy according to energy deposition in the detector. Systems may provide energy discrimination and material characterization based on low-energy and high-energy portions of incident X-rays. In a given energy region of medical CT, two physical processes dominate the x-ray attenuation: (1) Compton scatter and the (2) photoelectric effect. Thus, as known in the art, detected signals from two energy regions provide sufficient information to resolve energy dependence of the material being imaged and determine a relative composition of an object composed of two hypothetical materials.
A conventional third generation CT system may acquire projections sequentially at different peak kilovoltage (kVp) levels, which changes the peak and spectrum of energy of the incident photons comprising the emitted x-ray beams. Two scans are acquired—either (1) back-to-back sequentially in time where the scans include two rotations around the subject, or (2) interleaved as a function of the rotation angle requiring one rotation around the subject, in which the tube operates at, for instance, 80 kVp and 140 kVp potentials. When scanning sequentially, data obtained may be misregistered because of slight motion of the object between acquisitions. However, high frequency generators have made it possible to switch the tube voltage or potential of the X-ray source on alternating views. As a result, data for two energy sensitive scans may be obtained in a temporally interleaved fashion rather than in two separate scans made several seconds apart.
Thus, it is desirable to deliver microsecond or sub-microsecond current modulation of the electron beam and/or gridding in some imaging applications such that temporally interleaved scanning data may be obtained. Some technologies are capable of increasing or decreasing electron beam current, but such technologies achieve current modulation by changing an emitter temperature and thus the emitted beam current. Such current modulation processes are slow due to the thermal time constant of the emitter. That is, due to thermal mass of the filament it is not possible to achieve significant current modulation with this approach on a microsecond timescale.
To achieve a fast current response time, gridding technologies may be used to control electron beam operation electrostatically and modulate current, either via an intercepting or a non-intercepting grid. Typically, however, if high voltage is increased or decreased, the current will correspondingly increase or decrease as a consequence of respectively higher or lower electric fields at the emitter surface, which is a trend opposite that which is typically desired. That is, for an increased voltage it is typically desirable to have a decreased current, and vice versa. The higher current at lower voltage is desired to obtain sufficient X-ray flux at the detector surface, since the X-ray attenuation coefficient of materials decreases with increasing energy of the incident X-ray beam.
For a low tube voltage operation it is typically desirable to have a high current or tube mA, which in some applications is 1000 mA or greater at 80 kV, as an example. Correspondingly, it is typically desirable to have a low current or tube mA, 750 mA or less, at 140 kV. In today's tube, it is possible to achieve high emission by increasing filament temperature. However, as stated, this is a slow process and the temporal response of the temperature change of a filament is in the range of milliseconds. Second, increasing the temperature of the filament may curtail or limit filament life. Thus, for a fast kV switching operation, system operation and life requirements may limit the performance at desired current during low tube voltage operation of the fast kV switching operation.
Therefore, it would be desirable to have an apparatus and method capable of microsecond current modulation of an electron beam in an X-ray imaging device, while achieving high current emission without compromising emitter life.